The present invention relates to artificial joints, including those for use in hip and knee replacements.
There is a growing need for extension of the useful life of man-made prosthetic joints. For example, total hip replacement surgeries have been performed in the United States since 1971 with many hundreds of thousands of patients receiving needed artificial joints. Many of these installations have failed and have required more complicated and costly revisions or repairs. Many of the currently installed devices including hips and knees are expected to have a useful life of only 10-15 years depending on many factors including the level of activity of the recipients. A major factor contributing to these failures is the wear of one of the mated load-bearing surfaces which is most commonly and traditionally made from ultra-high molecular weight polyethylene (UHMWPE), a polymer with a microstructure having a crystalline phase material (usually present in the range of about 40 to 80% by volume) embedded in a noncrystalline phase matrix. Wear of this material results in loss of the original functionally engineered geometrically designed contours of artificial hips, namely, the wearing of a deep socket in the UHMWPE cup and related production of nanoscale particulate polymer wear debris. The debris, in turn, triggers an adverse foreign body response in the region of the installed device. This in vivo response can produce collateral damage leading to loosening of members of the device, subsequent dysfunction and pain and, ultimately, the need for surgical repair or revision.
Corrosion-resistant metal alloys and, less often, ceramics are used for the ball member of a hip joint which replaces the original anatomical femoral head. UHMWPE is the current material of choice for the cup member which replaces the surface of the pelvic opening known as the acetabulum. A great deal of research has been performed to find suitable and improved ball materials from the standpoint of reducing friction and wear against UHMWPE while assuring corrosion resistance and overall biocompatibility. These efforts carried out over the last 15-20 years have resulted in a short list of commonly accepted ball and stem materials for artificial hips and other prosthetic applications. This list includes AISI 316 stainless steels, cast cobalt-chromium-molybenum alloys, wrought cobalt-chromium alloys, unalloyed titanium, Ti-6Al-4V alloys, and cobalt-chromium-nickel alloys. Additionally, these alloys have been modified using various surface treatments for improved properties relative to prosthetic applications.
Another area of intense research on prosthetic materials has recently focused on improving the UHMWPE material used as a load bearing material opposing the metal or ceramic component of these artificial joint devices. This research has produced a variety of chemical and thermal processing methods which, taken in combination, alter the structure of the polymer. This structure-altered material is characterized in part by a reduction in scale and size of the crystalline phase lamellae within the microstructure as compared to the scale and size of lamellae in non-altered material.
Today, the cobalt-chromium-molybdenum alloys known under the standardization identity of ASTM F-75 and derivative identities are used most extensively for the ball material. ASTM F-75 type alloys contain up to 0.36 wt % carbon. This is primarily for the historical reason related to the adaptation and standardization of an alloy suitable for prosthetic applications from a cobalt-base, high temperature superalloy known to have corrosion resistance significantly exceeding that of stainless steel. This adaptation of the Haynes Stellite (HS-21), otherwise known as Modified Vitallium alloy, occurred at a time preceding the current more complete understanding of the mechanisms governing wear in prosthetic devices. There is a further historical relationship in the current use of this alloy as a prosthetic material in that the original composition developed by the Austenal Company was developed and used as a prosthodontic alloy in the late 1930""s. As is known in the metallurgical art, the simultaneous presence of carbon and chromium in a cobalt based alloy will result in the presence of hard phase chromium carbides dispersed within grains and at grain boundaries. Such a dispersion may be non uniform in cast structures with concentrations of carbides at grain boundaries. Because of the high atomic ratio of chromium to carbon in high chromium alloys, these carbides (Cr23C6) are more resistant to corrosion than the surrounding matrix. In the current art, it is known that in vivo corrosion of these cobalt alloys does take place as evidenced by recovery of metal ions in tissues and fluids surrounding installed prosthetic devices, and these carbides will, in time, end up in positive relief as asperities. Such asperities are believed to contribute to a ploughing effect in the accelerated wear of a matching UHMWPE cup surface of a prosthetic joint device. These carbides may also end up in relief in the manufacturing process because they inherently resist the polishing process used to produce the final surface finish of the femoral ball.
The evolutionary development of prosthodontic materials has resulted today in the availability of a wide range of noble metal alloys containing gold, platinum, palladium, silver, and small amounts of other elements used for control of processing, structure, and properties. These alloys have been proven to be corrosion resistant and biocompatible over long periods of time. In addition, they can be formulated and processed to produce hardness sufficient to withstand use in two body and three body friction and wear processes, in the chewing of food including hard foods and foods containing exogenous particles of very hard grit or foreign substances, such that the useful life of installed devices constructed from these noble metal alloys frequently exceeds 30 years or more. A unique characteristic related to these noble metal alloys is that they are not chemically capable of forming hard phases such as carbides, nitrides and oxynitrides as is well known in the metallurgical art and documented in the phase diagram literature. Thus, ploughing effects in wear produced by particles of this nature standing in microscopic relief do not occur with use of such noble metal alloys.
In the literature of the art on friction and wear (Friction and Wear of Materials, Rabinowitz, Wiley 1965), a quantitative law of adhesive wear (i.e. wear in the absence of third body abrasive particles) is given by the Holm equation (Holm, 1946). The worn-away volume, V, is stated as:
V=Kxc3x97Wxc3x97L/H,
where:
K is a constant depending on materials in contact and the extent of surface modification by lubricants, fluids, or other adsorbed chemical species,
W is the load,
L is the total sliding distance (for prosthetic devices in effect the useful life of the component most susceptible to wear, and
H is the hardness of the softer of the pair of materials in contact.
The hardness of the softer UHMWPE has certain published values taken at room temperature, most commonly 293xc2x0 K. This polymer is a crystalline polymer with a volume fraction of crystalline material ranging from about 40 to 80% depending on its thermal processing history. In vivo, this polymer operates at 310xc2x0 K when it is at equilibrium with its surroundings. The nominal crystalline melting point (Tmp) of this polymer is 410xc2x0 K. On an absolute temperature scale, the operating (or homologous) temperature of 310xc2x0 K is 0.76 Tmp or higher. In the field of materials science of crystalline solids, most commonly metals and ceramics, operation at a homologous temperature of 0.76 constitutes a high temperature use requiring appropriate temperature related design rules. UHMWPE is a low temperature material being applied in a high temperature regime when used in prosthetic devices.
In addition to the quantitative law of adhesive wear referred to above, an area of extreme interest to the early workers in the field of friction and lubrication of solids was the surface temperature produced during frictional rubbing of various pairs of materials, most commonly metals (Bowden and Tabor, The Friction and Lubrication of Solids, Oxford 1950). When one solid body slides over another, a significant portion of the mechanical work done against friction for devices where the coefficient is greater than zero is liberated as heat generated at or near the sliding surfaces. In artificial prosthetic joints installed in users (i.e. in vivo), this heat is dissipated by thermal conduction into the biological surroundings having a base reference temperature of 310xc2x0 K. Following known laws of heat transfer, the rate of thermal conduction of this frictional heat is controlled by a series of thermal resistances characterized by: 1. the dimensions and thermal conductivities of the materials of construction of the device, 2. the thermal contact resistances within the device in the cases of multipart or modular devices, 3. the thermal conductivities and thicknesses of cements or other substances used for device fixation, 4. the dimensions and thermal conductivity of bone into which the device is mounted or embedded, and 5. the thermal resistance or heat transfer film coefficients from bone and directly exposed device materials to the surrounding fluids and tissues around the device where such surroundings are fixed at an environmental temperature of 310xc2x0 K.
Embodiments of the present invention solve problems of the prior art with respect to maintaining load-bearing surfaces of prosthetic joints approximately at or below specified temperatures. Such thermal management solutions may yield extended life joints. Accordingly, in a first embodiment of the invention, a prosthetic joint includes a first member, with a first load-bearing surface and a second member, with a second load-bearing surface. The surfaces are slidingly disposed relative to each other defining a region of frictional contact during joint use. Heat is removed from the region during joint use by thermal conduction through at least one of the first and second members so that the first and second surfaces are maintained approximately at or below a specified temperature. The specified temperature may be less than a lowest melting point of a material of either of the surfaces. Alternatively, the specified temperature may be less than a temperature that destroys organic species indigenous to an in vivo joint. Also, the specified temperature may be equal to about that of surroundings in which the joint is used. The specified temperature may be equal to about 310xc2x0 K.
Thermal conduction may occur through the first member which may further have a first load-bearing portion that includes the first surface and a first body portion that is bonded to the first load-bearing portion. In another embodiment, the first member has greater than about seventy-five times higher conductivity than the second member when the conductivities are measured at about 310xc2x0 K. The first load-bearing portion may have a thermal conductivity higher than about 30 W/mxc2x0 K. The second member may contain polymeric material and may further have a second load-bearing portion that includes the second surface and a second body portion that is coupled to the second load-bearing portion. The polymeric material may be ultra-high molecular weight polyethylene and may have a thermal conductivity of less than about 0.4 W/mxc2x0 K. The prosthetic joint may serve as an artificial hip wherein the first member is ball-shaped and the second surface of the second member is cup-shaped and is sized to mate with the first member. In another embodiment, the second member may be ball-shaped and the first surface of the first member is cup-shaped and is sized to mate with the second member. The joint may also serve as an artificial knee.
In yet another embodiment, the first body portion also has a heat pipe core and a biocompatible casing surrounding the heat pipe core. The heat pipe core may have a thermal conductivity of greater than about 50 W/mxc2x0 K.
In a further embodiment, the first member also has a first load-bearing portion that includes the first surface, a first support portion that is bonded to the first load-bearing portion, and a stem that is mechanically coupled to the first support portion. The first support portion may further serve as a heat pipe in conducting heat from the first surface and may have a thermal conductivity of greater than about 50 W/mxc2x0 K. The stem may further have a heat pipe core and a biocompatible casing surrounding the heat pipe core. The biocompatible casing may be constructed from the same material as is the first-load bearing surface.
In another embodiment, the specified temperature may be less than a temperature at which the Meyer hardness of polyethylene decreases to about ninety percent of its Meyer hardness measured at about 310xc2x0 K. The first member may contain gold.
In yet another embodiment, a prosthetic joint has a first load-bearing surface, a second load-bearing surface, the surfaces slidingly disposed relative to each other defining a region of frictional contact, and a heat pipe coupled to at least one of the first and second surfaces. Heat is conducted via the heat pipe from the region to surroundings during joint use; the heat pipe having a thermal conductivity of greater than about 50 W/mxc2x0 K. One of the first and second load-bearing surfaces may contain gold. The heat pipe may be constructed from material selected from the group of molybdenum, molybdenum alloys, tungsten, and tungsten alloys.